Lung disease is the third largest cause of death in the United States of America, accounting for approximately one out of every seven adult deaths. It is estimated that 30 million Americans are living with chronic lung disease. Adult respiratory distress syndrome (ARDS) afflicts approximately 150,000 patients annually in the U.S. Despite advances in critical care, mortality remains around 40-50%.
Currently available therapies for patients with chronic respiratory failure include, for example, ventilation and extracorporeal membrane oxygenation (ECMO). Mechanical ventilation is effective for short-term support. Often, however, the excessive tidal volumes, airway pressure and oxygen fraction necessary to achieve sufficient gas exchange with mechanical ventilation can cause further damage to the lungs, creating ventilator-induced lung injury, including barotrauma, volutrauma and other iatrogenic injuries, further exacerbating acute respiratory insufficiency in many patients. ECMO systems are attractive since they closely simulate physiological gas exchange. However, they are complex in operation, can result in thrombosis, blood trauma, infection and bleeding due to the need for high levels of anti-coagulation, and limit patient mobility. Arteriovenous pumpless oxygenation may benefit some patients with acute lung failure but achievable flow is limited and its use is currently limited to intensive care units (ICU).
There have been efforts to develop more efficient and compact devices for respiratory and cardiopulmonary support systems. For example, attempts have been made to integrate multiple components of cardiopulmonary, ECMO systems into single structures so as to eliminate or minimize the need for the extension of lengthy, blood-filled tubes. Various integrated pump-oxygenators have been described (see, e.g., U.S. Pat. Nos. 5,217,689; 5,266,265; 5,270,005; 5,770,149; 4,975,247; 5,429,486; 6,963,222; and 6,730,267). There are drawbacks associated with these integrated pump-oxygenators, however. Such drawbacks include non-uniform blood flow through the fiber membranes and the existence of laminar boundary flow zones between blood cells and fiber membranes. Non-uniform blood flow across the fiber membranes results in hyper- and hypo-perfusion of the blood in flow paths. Hyper-perfusion does not grant any additional benefit once blood is oxygen-saturated, yet subjects the blood unnecessarily to prolonged exposure to artificial materials, thereby increasing risk of hemolysis and thrombosis. Hyper-perfusion occurs when oxygen-saturated blood is exposed to oxygenator fibers. The exposure to oxygenator fibers does not confer any benefit to the blood because it is already saturated with oxygen. Rather, the exposure unnecessarily increases shear stress and contact with synthetic material. Hypo-perfusion occurs when blood is incompletely saturated with oxygen before it is discharged from the oxygenator. In order to combat hypo-perfusion, longer flow paths and fiber membranes having larger surface areas (e.g. 2-4 m2) are used, resulting in extended contact of the blood with the fiber membrane surfaces, which, in turn, leads to blood activation thrombosis formation.
When the blood is passively pumped to flow through fiber membranes, a relatively thick blood boundary layer forms and impedes diffusion of oxygen to blood cells, which are not in direct contact with the filter membrane surface. Thus, the blood boundary layer significantly hinders gas transfer efficiency.
Therefore, gas exchange membrane surface areas of 2 to 4 m2 are typically required to provide the needed gas exchange for an adult. The bundle sizes utilized in open heart surgery are larger to compensate for the greater oxygen demand that occurs during patient re-warming. This is not critical in ECMO applications as the patient is maintained at normothermic conditions or only slightly less. Consequently, the volume of the devices is impractical for portable deployment. In addition, the pumps used in today's ECMO systems or proposed to be used in the above patents all can lead to device-induced blood trauma (hemolysis, and blood element activation) due to the technologies utilized (mechanical bearings, non-optimized flow paths and large prime volume).
Efforts to decrease the effect of the boundary layer include increasing shear rate and/or turbulence of the blood flow path by the introduction of secondary flows, for example, by directing blood to flow at a substantial angle, such as perpendicular, to the fiber membranes. U.S. Pat. No. 4,639,353, for example, discloses the use of an arrangement of bundles of hollow fibers perpendicular to the direction of blood flow via a series of flow guide structures. U.S. Pat. No. 5,263,924 discloses the use of an integrated centrifugal pump and membrane oxygenator comprising hollow fibers, which are displaced circumferentially in a ring around an impeller of the centrifugal pump and through which blood is pumped for oxygenation. Other efforts to decrease the effect of the boundary layer include actively rotating hollow fiber membranes or moving fiber membranes in the path of blood flow. The motion of membrane surfaces relative to the blood cells causes pumping and oxygenation to occur simultaneously and can disrupt the build-up of the boundary layer around the gas exchange surface. Examples of oxygenators with active gas exchange membranes include those described in U.S. Pat. Nos. 5,830,370; 6,723,284; and 6,503,450; and in the paper by Makarewics et al. (ASAIO 42: M615-619 (1996)).
Despite improvements in the performance and design of conventional blood pump-oxygenators, there remains a need for a more compact and efficient blood pump-oxygenator with less pronounced or fewer drawbacks. Thus, there is a need for a compact, efficient and non-traumatic pump-oxygenator with pumping function, low prime volume and low gas exchange membrane surface area.